U.S. patent application number 10/533118 was filed with the patent office on 2006-06-22 for balloon deployable stent and method of using the same.
Invention is credited to Vladimir Brailovski, Richard Gallo, Patrick Terriault.
Application Number | 20060136031 10/533118 |
Document ID | / |
Family ID | 32230362 |
Filed Date | 2006-06-22 |
United States Patent
Application |
20060136031 |
Kind Code |
A1 |
Gallo; Richard ; et
al. |
June 22, 2006 |
Balloon deployable stent and method of using the same
Abstract
The present invention provides a balloon-deployable stent having
a progressive expansion over time and a method for using such a
stent, thereby reducing restenosis. The stent has a progressive
radial expansion of an armature (12) comprising a material having
an elasticity allowing the self-deployment of the armature and of a
matrix (14) comprising a second material having a rigidity and a
conformation allowing a retention of the armature in a contracted
position. The stent is deployed with the help of a balloon
delivered into the armature, which allows an irreversible
deformation of the matrix during the inflation of the balloon and
enables a radial expansion of the armature.
Inventors: |
Gallo; Richard; (Montreal,
CA) ; Terriault; Patrick; (Verdun, CA) ;
Brailovski; Vladimir; (Montreal, CA) |
Correspondence
Address: |
Louis Tessier
P O Box 54029
Town of Mount-Royal
QC
H3P 3H4
CA
|
Family ID: |
32230362 |
Appl. No.: |
10/533118 |
Filed: |
October 29, 2003 |
PCT Filed: |
October 29, 2003 |
PCT NO: |
PCT/CA03/01676 |
371 Date: |
November 9, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60422489 |
Oct 31, 2002 |
|
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|
Current U.S.
Class: |
623/1.11 ;
623/1.15; 623/1.19 |
Current CPC
Class: |
A61F 2002/9511 20130101;
A61F 2002/91533 20130101; A61F 2210/0014 20130101; A61F 2/958
20130101; A61F 2230/0054 20130101; A61F 2250/0018 20130101; A61F
2002/91558 20130101; A61F 2/91 20130101; A61F 2/915 20130101 |
Class at
Publication: |
623/001.11 ;
623/001.15; 623/001.19 |
International
Class: |
A61F 2/84 20060101
A61F002/84; A61F 2/90 20060101 A61F002/90 |
Claims
1. A balloon-deployable and controlled radially expandable stent
comprising: an armature comprising a first material having an
elasticity allowing an expansion over time of said armature; a
matrix comprising a second material having a rigidity and a
conformation allowing a retention of said armature in a contracted
position; said stent being deployed with the help of a balloon
introduced into said armature, said balloon allowing an
irreversible deformation of said matrix during inflation of said
balloon and allowing expansion of the armature.
2. The stent of claim 1, wherein said armature and said matrix are
structures of similar rigidity.
3. The stent of anyone of claims 1 and 2, wherein said first
material is a shape memory alloy.
4. The stent of claim 3, wherein said metal shape memory alloy is
nitinol.
5. The stent of claim 1, wherein said second material is a polymer
and said deformation is plastic.
6. The stent according to anyone of claims 1 to 5, wherein said
matrix is fortified into a rigid geometry by rings.
7. The sent according to claim 6, wherein said rings are selected
in the group comprising a coating made of rings covering completely
said armature, rings braided around said armature, and rings
secured in slots provided on said armature.
8. The stent of anyone of claims 1 to 7, wherein said second
material has a rigidity of at least 1000 MPa, a yield strain below
about 8%, and an ultimate strain over about 100%.
9. The stent of anyone of claims 1 to 8, wherein said second
material is selected in the group comprising a polycarbonate and a
polyethylene.
10. The stent of anyone of claims 8 to 9, wherein said second
material further exhibits creep properties allowing a minimum loss
of 50% of an initial rigidity within 1000 hours.
11. The stent of anyone of claims 1 to 10, wherein the matrix
conformation is annular.
12. The stent of anyone of claims 1 to 11, further comprising a
retention sheath covering said matrix and said armature, and
recuperating expansion forces of said armature by preventing a
creep of said matrix.
13. A method of angioplasty in an artery of a patient comprising:
introducing and positioning in a vessel of the patient a
self-deploying stent having a progressive deployment comprising an
armature comprising a material having an elasticity allowing
self-deployment of the armature; and a matrix comprising a second
material having a rigidity and a conformation allowing a retention
of the armature in a contracted position; deploying the armature
using a balloon delivered in the armature, the balloon ensuring an
irreversible deformation of the matrix during inflation of the
balloon and allowing a self-deployment of the armature; and
removing the balloon from the vessel; whereby a progressive
self-deployment of the armature allows a positioning of the
armature at a predetermined position and a diminution of a risk of
restenosis.
14. The method of claim 13, wherein the armature comprises a shape
memory alloy.
15. The method of claim 14, wherein the shape memory alloy is
nitinol.
16. The method of anyone of claims 13 to 15, wherein the second
material is a polymer and wherein the deformation is plastic.
17. The method of claim 16, wherein the polymer has a rigidity of
at least 1000 MPa, a yield strain below about 8%, and an ultimate
strain over about 100%.
18. The method of anyone of claims 15 and 16, wherein the polymer
is selected in the group comprising a polycarbonate and a
polyethylene polymer.
19. The method of anyone of claims 17 and 18, wherein the polymer
further exhibits creep properties characterized by a loss of
rigidity of at least 50% of an initial rigidity thereof within 1000
hours.
20. The method of anyone of claims 13 to 19, further comprising
before step a) an expulsion of said stent from a retention sheath
covering the matrix and the armature and recuperating the expansion
forces of the armature by preventing a creep of the matrix.
21. The stent of claim 1, wherein said first material has
radio-opacity and rigidity properties comparable to metal.
Description
FIELD OF THE INVENTION
[0001] The present invention generally relates to stents. More
specifically, the present invention relates to a balloon deployable
stent and to a method making use thereof.
BACKGROUND OF THE INVENTION
[0002] Stents are typically used to enlarge or to liberate a
passageway in a vessel or a lumen.
[0003] For example, cardiovascular stents are used to increase a
diameter of a partially obstructed cardiovascular artery by forcing
an enlargement thereof through deployment of a metallic structure.
FIG. 1 illustrates the functioning mode of a cardiovascular stent,
as is well known in the art.
[0004] The installation of a cardiovascular stent is a
well-established technique in the art. More than 500,000
angioplasties per year are performed worldwide.
[0005] Two main materials are available in the marketplace for the
manufacture of cardiovascular stents: stainless steel and
nitinol.
[0006] In the case of stainless steel stents, an inflatable balloon
causes the deformation of the stent. The stent, in its contracted
state, is mounted on the balloon and introduced into the human
body. When the contracted stent mounted on the balloon is
positioned at a target location, the balloon is inflated, which
results in a plastic deformation of the stent. Next, the balloon is
deflated and pulled out of the artery, leaving the stent in a
deployed configuration against the walls of the artery. FIG. 2
illustrates such a stainless steel stent deployed and contracted on
an inflatable balloon.
[0007] Nitinol stents take advantage of an intrinsic property of
shape-memory alloy, whereby this material always regains an
original shape thereof if bent. The nitinol stent is introduced
into a catheter, which keeps the stent in a contracted position,
and moved in an artery to a target location. Once in position, the
stent is mechanically expulsed from the catheter and thereby
enabled to take its predetermined completely deployed
configuration, without the need for an inflatable balloon. Nitinol
stents are therefore self-deployable. FIG. 3 shows a nitinol stent,
which adopts its completely deployed form upon expulsion from the
catheter.
[0008] According to the "Handbook of Coronary Stents" (edited by P.
W. Serruys, Martin Dunitz Ltd, London, 1997), nitinol has been
employed since 1997 in a number of stents. In its superelastic
regime, nitinol is able to accommodate deformations in the order of
8% and tends to completely regain its initial non-deformed state.
In comparison, stainless steels such as the alloy 316L, which is
frequently used to manufacture stents, are able to accommodate a
reversible elastic deformation of about 0.1%. The elastic domain of
nitinol is approximately 80 times larger than that of conventional
metals like steel and aluminum. FIG. 5 schematizes the superelastic
behavior of nitinol, where E is the elasticity module;
.epsilon..sub.mf the strain at the end of the transformation;
.sigma..sub.ms the stress at the beginning of the transformation;
.sigma..sub.af the constraint at the end of the transformation; and
H the transformation module. Besides such a level of reversible
elastic deformations, nitinol has a high resistance comparable to
that of a metallic material, which allows an adequate dilation of
the artery and also guaranties stability over time.
[0009] A main concern is related to the fact that current
installation procedures of either the stainless steel or the
nitinol stents still cause a certain trauma of the vascular
walls.
[0010] For example, the pressure exerted by the inflatable balloon
for the installation of the stainless steel stents, so that the
latter espouses the inner walls of the vessel, traumatizes the
artery. One of the main problems associated with the use of
stainless steel stents as illustrated in FIG. 3 is that an
over-expansion is necessary during the inflation of the balloon to
compensate for an elastic springback effect. Indeed, when the
balloon is deflated, the stainless steel has a tendency to
contract, or to springback, due to an elastic component of the
total deformation. Consequently, to position the stent against the
wall of the artery in the best possible way, generally the stent
needs be inflated up to a diameter superior to that of the vessel
in order to compensate for the shrinkage or springback during
deflation. Such an over-expansion may damage the artery even
further and contribute to restenosis, while sub-expansion of a
stainless steel stent diminishes the interference constraint
between the stent and the walls, and may be detrimental since it
may be accompanied with increased rates of thrombosis and vessel
occlusion and, consequently, provoke the loss of stent-artery
contact. This effect is still more prominent if the artery relaxes
with time or if its diameter augments because of different
physiological reasons.
[0011] As far as nitinol stents are concerned, their diameter, once
completely deployed, may be greater than that of the artery. Hence,
during deployment, the nitinol stent is in contact with the artery
and the equilibrium of forces between the latter and the stent is
attained for a smaller diameter, thereby creating a permanent but
light pressure on the walls of the vessels. Indeed, due to the
particular behaviour of the nitinol stent, the stent keeps applying
a light pressure, which is practically constant while the diameter
of the artery increases, and continues to do so until the stent
attains its completely deployed diameter. Alternatively, should the
diameter of the artery decrease because of a spasm for example, the
stent offers a significant resistance to such a contraction.
However, the instantaneous liberation of a self-deploying nitinol
stent may also provoke an impact on the inner walls of a vessel
and, hence, causes trauma.
[0012] Traumas due to installation of the stents may contribute to
restenosis phenomenon (e.g. recurrence of vessel narrowing at the
site previously dilated). According to studies, about 30% of
angioplasties present a degree of restenosis within the first 6
months. As a result, a second intervention, for example an
introduction of another stent or a major brachytherapy operation
with the goal of effectuating bypass surgery, is needed. Restenosis
generates important costs for the healthcare system. It would
therefore be advantageous to provide a cardiovascular stent that
minimizes the trauma imposed on the coronary or peripheral vascular
systems during the deployment thereof.
[0013] It would therefore be advantageous to provide a balloon
deployable stent that does not necessitate an over-expansion during
placement overcomes the drawbacks associated with
sub-expansion.
[0014] Table 1 presents advantages and limitations associated with
stainless steel stents, such as illustrated in FIG. 2, and nitinol
stents, such as illustrated in FIGS. 3 and 4. TABLE-US-00001 TABLE
I Stainless Steel Stents Nitinol Stents (FIG. 2) (FIGS. 3, 4)
Advantages Limitations Advantages Limitations Possibility of
Necessary use Self-deploying Instantaneous gradually of inflatable
behaviour which deployment often applying balloon avoids the use of
affects its pressure on Over-inflation the inflatable positioning
and the walls of the may be balloon may traumatize arteries
necessary to Applying inner walls of because of the compensate for
permanent the vessel inflatable elastic pressure on walls, Elastic
behavior balloon springback hence, no elastic after its
Malleability Loss of springback installation, after pressure on the
which limits installation, inner walls of the interventions on
which is artery because of the secondary beneficial for the elastic
branches operations on springback the secondary branches
[0015] In summary, stainless steel stents satisfy the gradual
deployment property. Their installation method with an inflatable
balloon allows a precise and gradual positioning. However, they
generally require an over-deployment and often suffer from elastic
springback. Nitinol stents, on the other hand, may adapt to
variations of the vessel diameter. Nevertheless, their
self-deploying capacity and abrupt deployment may compromise their
positioning.
[0016] A stainless steel stent requires an inflatable balloon to
deform, and once deformed, it does not tend to regain its initial
contracted state, whereas a nitinol self-deployable stent always
tends to regain its completely deployed state, without the need for
an inflatable balloon. Indeed, as shown in FIG. 3, when the nitinol
stent is expulsed from the catheter, which keeps it in a contracted
position, it deploys itself instantly to come back to its initial
state. Lastly, this capacity to accommodate a great deformation
facilitates the progression of the stent through the often tortuous
vessels (e.g. arteries and other lumens) of the human body.
[0017] Therefore, there is still a need for a stent, which
gradually deploys, allowing a precise and controlled installation
while avoiding an abrupt mechanical action, and which, once
deployed, exerts a continuous pressure on the walls of the artery
or vessel even if a diameter thereof increases, through a
controlled radial expansion, thereby minimizing the downside of an
elastic springback effect.
OBJECT OF THE INVENTION
[0018] The present invention therefore relates to an improved
balloon deployable stent and to a method making use thereof.
SUMMARY OF THE PRESENT INVENTION
[0019] The present invention provides a balloon-deployable and
controlled radially expandable stent comprising an armature
comprising a first material having an elasticity allowing an
expansion over time of the armature; a matrix comprising a second
material having a rigidity and a conformation allowing a retention
of said armature in a contracted position; the stent being deployed
with the help of a balloon introduced into the armature, the
balloon allowing an irreversible deformation of said matrix during
inflation of the balloon and allowing expansion of the
armature.
[0020] The invention further provides a method of angioplasty in an
artery of a patient comprising: introducing and positioning in a
vessel of the patient a self-deploying stent having a progressive
deployment comprising an armature comprising a material having an
elasticity allowing self-deployment of the armature; and a matrix
comprising a second material having a rigidity and a conformation
allowing a retention of the armature in a contracted position;
deploying the armature using a balloon delivered in the armature,
the balloon ensuring an irreversible deformation of the matrix
during inflation of the balloon and allowing a self-deployment of
the armature; and removing the balloon from the vessel; whereby a
progressive self-deployment of the armature allows a positioning of
the armature at a predetermined position and a diminution of a risk
of restenosis.
[0021] Other objects, advantages and features of the present
invention will become more apparent upon reading of the following
non-restrictive description of embodiments thereof, given by way of
example only with reference to the accompanying drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
[0022] FIG. 1, which is labeled "prior art", illustrates the way by
which a cardiovascular stent according to the art is inserted into
a partially blocked artery (A); the deployment of the stent (B);
the enlarged artery (C);
[0023] FIG. 2, which is labeled "prior art", shows a stainless
steel stent according to the art deployed (upper section) and
contracted on an inflatable balloon (bottom section);
[0024] FIG. 3, which is labeled "prior art", illustrates a nitinol
stent according to the art covered by a catheter (upper section);
starting to self-deploy as the catheter is withdrawn (towards the
left, in the middle section); and completely deployed after
expulsion of the catheter (bottom section);
[0025] FIG. 4, which is labeled "prior art", illustrates a) a
Radius.TM. stent comprising five zigzag segments; b) a photo of the
stent taken with a scanning microscope demonstrating the precision
by a laser cut;
[0026] FIG. 5, which is labeled "prior art", illustrates the
superelastic behavior of nitinol;
[0027] FIG. 6 illustrate embodiments of a polymeric matrix on the
metallic armature according to the present invention;
[0028] FIG. 7 schematizes a stress-strain relationship employed in
the calculations for the design of the polymeric matrix of a stent
according to the present invention
[0029] FIG. 8 is a graphic representation of the armature's
external diameter evolution;
[0030] FIG. 9 illustrates an embodiment of a polymeric ring: A)
front view; B) lateral view;
[0031] FIG. 10 illustrates polymeric rings braided over-under
around the armature according to an embodiment of the present
invention;
[0032] FIG. 11 illustrates polymeric rings secured into slots
provided on the armature according to a further embodiment of the
present invention;
[0033] FIG. 12 illustrates graphically, the functioning of the
stent comprising the nitinol armature and of the polymeric rings,
in which A corresponds to the completely deployed stent; B
corresponds to the contracted armature; C corresponds to the
non-deformed rings; D corresponds to the established equilibrium
between the armature and the rings (passage from C to D), the
armature passing from B to D; the deployment of the balloon brings
the metallic armature from point D to point E and the rings from
points D to F. The elastic springback of the polymeric rings is
illustrated by the passage from point F to point G, while the
armature passes from point E to point G; and
[0034] FIG. 13 illustrates graphically the variation in rigidity of
the ring caused by the creep property; I corresponds to the
rotation of the slope in relation to the pivot; H corresponds to
the new equilibrium position.
DETAILED DESCRIPTION OF THE INVENTION
[0035] The present invention provides a balloon deployable stent,
which has the property of having a progressive radial expansion
over time, and a method making use thereof.
[0036] According to a first aspect of the present invention, as
illustrated in FIG. 6, a balloon deployable stent according to the
present invention comprises an armature 12 and a matrix 14.
[0037] The matrix 14 is mounted, either glued or affixed for
example, to the armature 12 in a contracted state thereof.
[0038] The armature 12 is made of a first material, which may be
selected in order to obtain radio-opaque and rigid properties in
the artery, which may prove to be interesting properties in the
course of an angioplasty intervention, for example, which are
comparable to those of metal.
[0039] The matrix 14 is made in a second material, which may be
selected between materials which, in time, gradually lose their
mechanical properties, thereby allowing a gradual and controlled
expansion of the armature 12.
[0040] The stent of the present invention may further comprise a
retention sheath made of a third material. Non-limiting examples of
this retention sheath include a sheath per se, as well as a
polymer, collagen-like or biological glue material, which inhibits
the expansion of the armature 12 and of the matrix 14.
[0041] It is to be understood that the term "material" as used
herein for the armature 12, matrix 14, sheath or other element of a
stent according to the present invention refers to at least one
material, and thus covers the combination of many materials, e.g.
an alloy, a mix of polymers, a mixture, etc. Obviously, when
mixtures of materials are used, the properties of the mixtures
satisfy the requirements associated with the particular uses of the
stent. For example, when a mixture of materials is used for the
matrix 14, at least one of the materials in the mixture loses its
mechanical properties in time, thereby enabling a gradual expansion
of the armature 12 over time.
[0042] The armature 12 may be made in a metal such as nitinol,
which is a shape memory alloy comprising titanium and nickel. In
this case, the nitinol armature 12 coupled to the matrix 14 is
installed on an inflatable balloon to facilitate the installation.
The matrix 14, which tends to loose its mechanical properties,
allows securing the stent maintained in place on the balloon
thereby delaying the deployment thereof. When the balloon is
inflated, it deforms the matrix 14 in an irreversible fashion.
Having lost its mechanical property, the matrix 14 then no longer
restricts the self-deployment of the nitinol armature 12.
[0043] It is to be noted that the matrix 14 may be selected in such
a way that the loss of its mechanical properties occurs at a
relatively low temperature, such as the temperature of the human
body (370C).
[0044] Furthermore, creep properties of the material(s) of the
matrix 14 may be put to use in order to delay the deployment of the
stent after the positioning of the stent. It may be briefly
reminded that, when a mass is suspended at the end of a wire
constituted from a certain material, an instant elongation
proportional to the suspended mass may be observed. This elongation
is the manifestation of the material's elastic behavior. If this
mass is left in place for a sufficient amount of time, a
progressive increase in the elongation of the wire over time may be
recorded. This progressive elongation is the manifestation of what
is referred to as the creep phenomenon. The additional lengthening
of the wire is thus an indicator that the material weakens under
the creep. For metallic materials, creep generally takes place at
high temperatures, (e.g. in the order of several hundreds degrees
Celsius). In the present invention, a material exhibiting creep at
human body temperature may be selected.
[0045] The matrix 14 of the progressive radially expandable stent
of the present invention may comprise at least in part polymeric
materials, which may creep at low temperatures. In selecting the
polymer to make the matrix 14, the following properties are
contemplated:
[0046] 1) a sufficient rigidity allows maintaining the stent in a
contracted position on the balloon. During the storage of the
stent, a retention sheath (or catheter) may be added, e.g. a hollow
cylinder in which for example the stent, including the matrix,
mounted on the balloon, may be introduced, so as to avoid an
unnecessary creep of the matrix during the storage of the stent.
The stent may be maintained in a contracted position when not in
the retention sheath, providing the matrix is sufficiently rigid. A
polymer with a module of sufficient rigidity may then be selected
in order to keep the deployment of the armature;
[0047] 2) a capacity of plastic deformation of the polymer during
dilation caused by the inflation of the balloon, i.e. without
elastic return (or possibly with a negligible elastic return), may
allow to avoid or minimize a return of the armature towards its
contracted position. Therefore, a polymer with a low yield strains
(passage from the elastic to the plastic regime), relative to
certain silicones from which the reversible elastic strains are
several hundred percents, may be advantageous;
[0048] 3) a capacity of plastic deformation of the polymer during
the inflation of the balloon, without tendency of rupturing or
fissuring of the polymeric matrix, which might liberate particles
in the vessel (e.g. in the blood) may be obtained providing a
sufficiently high ultimate strain. It is believed to be within the
reach of a person of ordinary skill in the art to select the
materials for the armature and matrix so as to be compatible with
their clinical use;
4) a capacity to creep at human body temperature under the forces
generated by the armature may further contribute to a progressive
deployment over time;
[0049] 5) a biocompatibility and conformance with USP standards is
required so that the stent may be used into the human body as an
implant. Non-restrictive examples of the different classes of
polymer used for medical purposes and which satisfy the USP
standards comprise certain types of Urethane-polycarbonate
(Bionatel for example), of polycarbonate (Makrolonl for example),
of polyethylene or of polypropylene (from Huntsman or Montell for
example).
[0050] As a way of example, the polymer Makrolonl Rx 2530 (Bayer)
has proven to satisfy the previously enunciated features. A high
density polyethylene like DMDA-8920 Natural 7 provided by
Petromont, and characterized by a rigidity of about 1000 MPA, a
yield strain of about 7%, an ultimate strain of over 400% and creep
properties whereby about 90% of the initial rigidity is lost after
100 hours at 37.degree. C. under a stress of 8.7 MPA, may also be
contemplated for example.
[0051] Generally stated, the matrix of the present invention may be
made of a polymer having a high rigidity of at least 1000 Mpa, a
low yield strain below about 8%, a large ultimate strain over about
100%, and creep properties allowing a minimum loss of 50% of the
initial rigidity.
[0052] It is to be noted that alternative materials, which loose
their mechanical properties over time, may be used to form the
matrix, such as materials subject to biodegradation for
example.
[0053] Therefore, according to a second aspect of the present
invention, a method for angioplasty comprises introducing and
positioning in a vessel of the patient a self-deploying stent
having a progressive deployment comprising an armature comprising
of a material having an elasticity allowing self-deployment of said
armature and a matrix comprising a second material having a
rigidity and a conformation allowing a retention of the armature in
a contracted position; deploying the armature using a balloon
delivered in the armature, the balloon ensuring an irreversible
deformation of the matrix during inflation of the balloon and
allowing a self-deployment of the armature; and removing the
balloon from the vessel; whereby a progressive self-deployment of
the armature allows a positioning of the armature at a
predetermined position and a diminution of the risk of
restenosis.
[0054] Numerical analyses were carried out in order to size the
polymeric matrix 14 to be added to the metallic armature 12 and to
verify if, theoretically, the armature-polymer unit may pursue a
progressive and retarded deployment after having been dilated by
the inflatable balloon.
[0055] Rather than attempting to size the polymeric matrix 14 by
considering simultaneously every desired function, such as initial
rigidity, creep property, for example, first calculations were
carried out considering only the mechanical behavior of the polymer
so as to meet the two following requirements: i) keep the armature
12 in a contracted position before the deployment of the inflatable
balloon, and ii) deform itself irreversibly, i.e. by plastic
deformations, during the inflation of the balloon without reaching
the ultimate strain.
[0056] The creep behavior of the polymer will be verified to ensure
that the stent offers an additional progressive deployment after
installation
[0057] The stent used for the numerical validation is a nitinol a
self-deploying Radius.TM. of SciMED, fabricated from a tube cut to
a desired geometry with a laser. This geometry comprises a number
of zigzag segments linked to each other by three bridges, as
illustrated in FIG. 4, drawn from the Handbook of Coronary Stents,
supra.
[0058] The dimensions used for the numerical validation are
presented in Table 2. TABLE-US-00002 TABLE 2 Nominal diameter 3.0
mm External diameter completely deployed 3.75 mm Width of a zigzag
segment 2.5 mm Number of zigzags per segment 9 Thickness of the
stent (tube wall) 0.11 mm Width of the strut 0.11 mm
[0059] It should be noted that even if the diameter of the
completely deployed nitinol armature is 3.75 mm, the company
indicates that it is a stent of 3.0 mm. Therefore, it is
recommended to use a stent of 3.75 mm for use into an artery of
about 3.0 mm.
[0060] The properties of nitinol depend greatly on the alloy
composition and of the thermal treatment it has received. Average
properties as derived from the literature are used for the
numerical simulations. The stress-strain relationship, also called
the material law, used to stimulate the superelastic behavior of
nitinol and the value of the parameters describing this behavior
are respectively shown and given in FIG. 5 and Table 3.
TABLE-US-00003 TABLE 3 Elasticity modulus (E) 80 000 MPa Poisson
coefficient (.upsilon.) 0.3 Strain at the end of transformation
A.fwdarw.M (.epsilon..sub.mf) 8.0% Strain at the start of
transformation A.fwdarw.M (.sigma..sub.ms) 500 MPa Stress at the
end of transformation M.fwdarw.A (.sigma..sub.af) 250 MPa
Transformation module (H) 2500 MPa
[0061] The bilinear material law representing the superelastic
behavior of nitinol is used to compute the response of a
tridimensional nitinol structure. An hysteresis in the material
law, whereby the path followed during unloading is not the same as
that followed during loading, is observed. This property of the
material will transpose itself to the behavior of the nitinol
stent. In fact, it is observed during the numerical simulations
that the response in contraction of the metallic armature does not
follow the same path at the time of its deployment.
[0062] During the deployment of the stent by the inflatable
balloon, the polymer's behavior may be characterized as being
"elasto-plastic". Until a certain stress value, referred to as the
yield stress .sigma..gamma. is reached, the material behaves
elastically without manifesting a plastic deformation. The initial
rigidity of the material is given by the Young's modulus (E). When
the stress exceeds the yield value, plastic strains are induced in
the material and a residual deformation is thus obtained if the
stress is subsequently taken back to zero. This plastic regime may
be observed until rupture, that is when the deformation reaches a
value .epsilon..sub.B referred to as the ultimate strain. The
rigidity at the time of plastification is largely inferior to that
obtained during the elastic behavior.
[0063] FIG. 7 schematizes a stress-strain relationship used in the
calculations to model the polymeric matrix's behavior. The values
of the different parameters used for the calculations are given in
Table 4, as taken from the technical data for Bayers Makrolon Rx
2530. TABLE-US-00004 TABLE 4 Elasticity modulus (E) 2400 MPa
Poisson coefficient (.upsilon.) 0.3 Yield stress (S.sub.Y) 65 MPa
Ultimate stress (S.sub.B) 75 MPa Yield strain (.epsilon..sub.Y) 6%
Ultimate strain (.epsilon..sub.B) 120%
[0064] To calculate the behavior of the nitinol metallic armature,
a finite element method may be employed, considering only half of a
zigzag is considered, since during a uniform contraction, the nine
zigzags form a cylindrical segment that is subjected to the same
efforts and deformations. Therefore, it is not necessary to repeat
the same calculations many times. A finite element mesh used for
the analysis is constituted of 893 nodes and 2,955 elements in a
tetrahedral form (pyramid with a triangular base). All these
elements follow the law of nitinol's bilinear behavior described
hereinbefore. To cause the deformation, punctual forces F are
applied at the extremities of the semi-zigzag. Therefore, an
increase in these forces yields a contraction of the stent. To take
into account the fact that the real geometry is not only
constituted by one semi-zigzag, but by a number of them,
geometrical restrictions are introduced as "boundary conditions".
In the present case, these boundary conditions reflect the fact
that the symmetry surfaces always lie on the same plane during the
structure deformation.
[0065] By increasing the intensity of the forces, the stent
contracts itself and the decrease in diameter of the stent in
relation to the applied force may then be calculated. The graph in
FIG. 8 shows the evolution of the external diameter of the
armature. The diameter has a value of 3.75 mm when the stent is in
the completely deployed position. An hysteresis related to the fact
that the stent follows a different path during the contraction and
during the progressive deployment is clearly noticeable. The
hysteresis is intrinsically considered in the bilinear material
law, which simulates the super-elastic behavior of nitinol.
[0066] In addition to calculating the stent contraction in relation
to the applied force, the finite element analysis allows estimating
the stresses in the material. A stress rise near the extremities of
the semi-zigzag is thus observed, while a central part thereof
appears to be practically unsolicited. Maximal stresses generated
in the structure are approximately 600 MPa, which corresponds to
less than 5% of strain. Because nitinol is able to accommodate
close to 8% of strain, it may then be concluded that this stent may
be contracted to levels equivalent to those reached during the
analysis without getting damaged and even further.
[0067] Returning to FIG. 6 of the appended drawings, clearly the
addition of a polymeric matrix on the armature may be done in
several ways adaptable by a person of ordinary skill. For example,
as shown in FIG. 6A, the polymer may completely cover the metallic
armature with a layer thereof. However, this solution appears
technically inconvenient, since the metallic armature and the
polymeric matrix would have more or less the same geometry, while
the nitinol is approximately 75 times more rigid than Makrolon
(rigidity module of 80 000 MPa for nitinol, compared to 1100 MPa
for Makrolon). Therefore, it results that the layer of polymer
would not be able to maintain the metallic armature in a contracted
position.
[0068] Therefore, since the geometry is similar, the rigidity of
the materials is roughly the same. Moreover, addition of this layer
of polymer is intended to be performed when the stent is completely
contracted. From the point of view of fabrication, mechanical means
used to maintain the metallic armature in a contracted position may
prevent the application of a uniform layer of polymer.
[0069] As FIG. 6B shows, a possible solution to obtain two
structures of similar rigidity, the structures being the polymeric
matrix and the metallic armature, comprises using rings of polymer.
Since the polymer is clearly less rigid than nitinol, structures
with a similar rigidity may be obtained by fortifying the polymeric
matrix into a rigid geometry. Therefore, the nitinol armature is
materially rigid and structurally flexible, while the polymeric
ring is materially flexible and structurally rigid.
[0070] It is to be noted that the rings of polymer may be
positioned in a number of alternative ways. For example, a complete
coating 14 covering completely the armature 12 may be used as
illustrated in FIG. 6 A, or the rings 14 may be braided around the
armature 12 (see FIG. 10), or secured in slots 16 provided on the
armature 12 (see FIG. 11).
[0071] Dimensions of the polymeric ring-like structure yielding a
desired global rigidity may be calculated. The present analysis is
based on the interior diameter of the ring (1.89 mm), which
corresponds to the exterior diameter of the metallic armature for
the maximal contraction reached during the previous analysis. The
other dimensions, being the thickness of the wall at 0.025 mm and
the width of the ring at 0.05 mm, are determined by applying a
calculation algorithm (FIG. 12). Finally, the material of the ring
is exemplified here with Makrolon, the properties of which have
been discussed hereinbefore. Of course, a person of ordinary skill
will understand that the present invention may use alternative
material in the polymer matrix.
[0072] For the calculation of the behavior of the ring, it may be
assumed that the stress in a section is uniform. In fact, because
of the geometry of the ring, its rigidity in flexion, referred to
as radial crushing, is negligible in comparison to its rigidity in
traction caused by the uniform dilation, i.e. increase in
circumference. Therefore, the ring has a tendency to take the form
of a polygon during its dilation by the inflatable balloon. When
assuming that the ring dilates into a polygon form, the stress in
the section of the ring may thus be considered simply as axial
traction. Moreover, it is assumed that the ring is deformed by the
reaction forces of the metallic armature on the ring, which are
thus the equivalent of the forces F considered hereinabove to
deform the nitinol armature.
[0073] The study of the behavior of the ring consists in evaluating
the force F required to increase the diameter of the ring from an
initial value .phi..sub.0 to a given value .phi.. Mathematical
relations applicable to a regular polygon enable us to link the
length of a side of the polygon S to its radius R
[0074] When the ring is deformed, the length S of the side of the
polygon increases, and this from S.sub.0. The imposed stress to the
ring may then be calculated by the definition of engineering
strain, that is, the ratio of the lengthening on the initial
length: .epsilon.=.DELTA.L/L.sub.0=(S-S.sub.0)/S.sub.0
[0075] The last notion that must be presented before presenting the
calculation algorithm concerns the equilibrium of the forces. The
force F must be balanced by the internal stresses a generated in
the polymeric ring. However, it was discussed before that these
stresses are supposed to be constant and normal at the section of
the ring. Therefore, it is possible to write the following
equilibrium equation where b and t are respectively the thickness
and the width of the polymeric ring: F=2.sigma.b t
sin(20.degree.)
[0076] By using the results presented previously, the following
method may be used to estimate the behavior of a polymeric ring.
The goal is to develop a method enabling to link the applied force
F on the ring in relation to the diameter .phi. of the ring.
1) Impose a value to the diameter of the deformed ring .phi.
2) Calculate the length of the side of the polygon: S=0.342
.phi.
3) Calculate the stress in the ring by using the initial length
S.sub.0=0.646 mm: .epsilon.=(S-S.sub.0)/S.sub.0 4) For a given
strain value .epsilon., deduct the stress .sigma. with the help of
the material law for Makrolon; 5) Calculate the force necessary to
generate this stress with the help of the equilibrium equation F=2
.sigma.b t sin (20.degree.) when b=0.05 mm and t=0.025 mm:
F=8.5510.sup.4 .sigma.
[0077] In summary, starting with a given value for the interior
diameter .phi. of the polymeric ring, it is possible to estimate
the necessary force F to reach this diameter. Table 5 gives the
results of the calculations that were made with the help of the
algorithm using the 5 steps presented above. The diameter of the
ring is dilated from the initial internal diameter of 1.89 mm to a
diameter of 3.10 mm and then, the force F is pulled back to 0.
TABLE-US-00005 TABLE 5 .PHI. (mm) S (mm) .epsilon. .sigma. (MPa) F
(N) 1.89 0.646 0.000 0.0 0.000 1.92 0.657 0.016 38.5 0.033 1.95
0.667 0.032 58.5 0.050 2.00 0.684 0.059 64.9 0.055 2.05 0.701 0.085
66.7 0.057 2.10 0.718 0.112 67.6 0.058 2.15 0.735 0.138 68.3 0.058
2.20 0.752 0.165 68.9 0.059 2.25 0.770 0.191 69.4 0.059 2.30 0.787
0.218 70.0 0.060 2.35 0.804 0.244 70.3 0.060 2.40 0.821 0.271 70.6
0.060 2.45 0.838 0.297 70.9 0.061 2.50 0.855 0.324 71.2 0.061 2.55
0.872 0.350 71.4 0.061 2.60 0.889 0.376 71.6 0.061 2.65 0.906 0.403
71.7 0.061 2.70 0.923 0.429 71.9 0.061 2.75 0.940 0.456 72.0 0.062
2.80 0.958 0.482 72.2 0.062 2.85 0.975 0.509 72.3 0.062 2.90 0.992
0.535 72.5 0.062 2.95 1.009 0.562 72.6 0.062 3.00 1.026 0.588 72.7
0.062 3.05 1.043 0.615 72.9 0.062 3.10 1.060 0.641 73.0 0.062 3.07
1.050 0.626 36.5 0.031 3.04 1.041 0.611 0.0 0.000
[0078] Table 5 stresses the maximal strain reached during the
analysis, that is 64.1% of strain. However, this value is largely
inferior to the ultimate strain of 120% from which the fissuring or
the rupture of the Makrolon may arise. The chosen conception is
thus safe in relation to that aspect.
[0079] The non-linear behavior may be observed by considering the
force applied on the ring in relation to the diameter .phi.
thereof. It appears that the ring acts more or less like an elastic
when the dilation force is inferior to 0.05 N. Beyond this value,
plastification occurs and the ring deforms itself considerably
under the effect of a very light increase in the applied force.
Moreover, an elastic return (elastic springback) occurs when the
force on the ring is completely released after having been dilated
to a diameter of 3.10 mm. A residual deformation is also observed,
because the ring does not return to its initial state with a
diameter of 1.89 mm, but rather to a dilated state with a diameter
of 3.04 mm. This irreversible behavior is the consequence of the
elasto-plastic characteristic of the polymer.
[0080] The analysis carried-out previously may now be combined to
determine the simultaneous behavior of the nitinol armature and the
polymeric rings. A working model comprises two rings per zigzag
segment of the metallic armature. If, for example, a metallic
armature comprises 7 zigzag segments, 14 such rings are assumed on
the metallic armature, the internal diameter of the rings being the
same as the external diameter of the armature, and referred to as
the interface diameter. Of course, it will be realized that
depending on the polymer used, the shape and design of the
ring-like members as well as the number of such ring-like members
etc. may be varied.
[0081] FIG. 12 allows understanding the functioning of the stent
comprising the nitinol armature and the polymeric rings. In this
graphic, the nitinol armature is firstly completely deployed, as
represented by point A in the graphic, and the diameter of the
armature is 3.75 mm. The armature is then contracted to a diameter
of 1.89 mm (point B), under a force F of 0.06 N. At that time, the
non-deformed polymeric rings, which also have an internal diameter
of 1.89 mm, are introduced around the metallic armature. The
non-deformed rings correspond to point C on the graph. Then, the
contraction force applied on the metallic armature is completely
released, thereby reestablishing the equilibrium between the
armature and the rings, which implies that the rings pass from
points C to D, while the armature passes from points B to D. Point
D represents the equilibrium state before the deployment commanded
by an inflated balloon. The diameter of the structure is 2.0 mm. As
observable in FIG. 12, such a design of the polymeric rings allows
retaining the nitinol metallic armature in a contracted
position.
[0082] The deployment of the stent in the artery is performed using
an inflatable balloon. The interface diameter is then increased
from the equilibrium position (2.0 mm) to a value of 3.1 mm. Thus
the metallic armature passes from point D to point E, while the
rings do the same from point D to point F. The balloon is then
deflated and removed, which results in the armature and the rings
reaching a second equilibrium position. A light contraction due to
the elastic springback of the polymeric rings that pass from points
F to G is observed, while the armature passes from points E to G
according to a path describing an hysteresis (curved arrow,
starting at point E). Point G represents the equilibrium position
after the inflatable balloon has been deflated and removed from the
artery. The interface diameter of the stent is then 3.07 mm.
Clearly, in this example, the rings appear to sustain a great
irreversible deformations during the inflation of the balloon.
[0083] At point G, the stent comprising the nitinol armature
reinforced by the Makrolon rings is installed in the artery. The
equilibrium forces are then approximately 0.032 N. According to
computations, these forces generate stresses of about 37 MPa in the
polymeric rings. According to the Makrolon 3100 creep data
enumerated above, a loss in rigidity of the material of about 33%
is observed when the material is subjected to stresses of some
thousands psi (which represents about ten or so MPa) at a
temperature neighboring that of the human body (104.degree. F.) for
a time lapse of 1000 hours (42 days). At point G, the rigidity of
the ring is 1.03 N/mm, as determined from the slope of the curve. A
rigidity of 0.68 N/mm may thus be set forward after 1000 hours at
37.degree. C. FIG. 13 shows the variation in rigidity of the ring
caused by the creep phenomenon. The loss of rigidity may be
visualized graphically by the rotation of the slope in relation to
the pivot 1. The new equilibrium position then becomes point H,
which represents the intersection of the new rigidity of the ring
with the curve modeling the increase of the diameter of the
armature. Point H indicates a diameter of 3.09 mm after 1000 hours
of creep, which represents an increase of 0.02 mm in relation to
the diameter before creep.
[0084] The increase in diameter due to creep as exemplified
hereinabove is rather weak. However, analyses demonstrate that it
is possible to increase the diameter of a stent on a long period of
time following the surgical intervention. The material exemplified
here does not creep sufficiently to permit a desired increase of
the diameter of the stent over a prolonged period of time following
implantation. A loss of more than 90% of the rigidity of the
polymeric material may allow a post-operation deployment of about
0.2 or 0.3 mm, which would come closer to the sought-out
performances. Therefore, the use of a polymeric material having
similar mechanical properties as the Makrolon exemplified herein,
while being able to creep more, would enable the reaching of the
objectives of the most preferred conceptions. For example,
polyethylene such as DMDA-8920 polyethylene, which may have a loss
of more than 90% of the rigidity, may allow a post-operation
deployment of the stent.
[0085] The results mentioned above demonstrate the feasibility of a
stent comprising a nitinol armature and polymeric bands. This stent
is able to deploy itself on a long period of time (from several
weeks to several months) in an independent and progressive fashion,
once the surgical operation is completed. Moreover, the
installation of the stent in the artery is performed using an
inflatable balloon, which allows a control thereof.
[0086] Dimensioning of the different components of the stent as
determined is then suggested. It is shown that the polymeric rings
are important structural elements of the stent since they delay the
deployment of the nitinol armature. It is shown for example that
rings of 0.025 mm of thickness and 0.050 mm in width may assure a
post-operation deployment of a Radius.TM. nitinol stent of 3.75 mm.
According to the dimensions mentioned, these rings have the
necessary rigidity to retain the metallic armature in a contracted
position before the deployment by the inflatable balloon, while
simultaneously still offering great irreversible deformations
during the inflation of the balloon. The efficiency of the creep in
allowing a retarded and progressive deployment may be tested
experimentally in animals.
[0087] From the foregoing, it should now be apparent that the
present invention is a cardiovascular balloon-deployable stent
having a controlled radial expansion over time enabling its precise
installation.
[0088] The stent of the present invention minimizes the damage to
the vessel, which may be associated with an abrupt mechanical
expansion.
[0089] Moreover, the stent of the present invention minimizes the
springback elastic effect in view of its continuous pressure on the
artery walls for a prolonged period of time, after the withdrawal
of the balloon that has served for its initial deployment.
[0090] Other numerical analyses may further be realized to take
into account the behavior of the artery during the deployment of
the stent.
[0091] In addition to characterizing the creep properties of the
polymeric material used, the analyses in animals may help to verify
experimentally the global behavior of the stent. Prototypes may be
made and tested.
[0092] As used herein, the term "cardiovascular" encompasses the
coronary and peripheral vascular systems. Hence, the stents of the
present invention are not limited to a use in coronary angioplasty.
Indeed, the stents of the present invention find use in any disease
or condition in which a stenting of a vessel or lumen would be
beneficial.
[0093] Although the present invention has been described
hereinabove by way of possible embodiments thereof, it may be
modified without departing from the nature and teachings thereof as
defined in the appended claims.
* * * * *